This invention relates generally to magnetic resonance imaging (MRI) systems and, more particularly, to radio-frequency (RF) coils in such systems.
Magnetic Resonance Imaging (MRI) utilizes hydrogen nuclear spins of the water molecules in the human body, which are polarized by a strong, uniform, static magnetic field of a magnet (typically denoted as B0—the main magnetic field in MRI physics). The magnetically polarized nuclear spins generate magnetic moments in the human body. The magnetic moments point or are aligned parallel to the direction of the main magnetic field B0 in a steady state and produce no useful information if they are not disturbed by any excitation.
The generation of nuclear magnetic resonance (NMR) signals for MRI data acquisition is accomplished by exciting the magnetic moments with a uniform radio-frequency (RF) magnetic field (typically referred to as the B1 field or the excitation field), for example, by applying a uniform RF magnetic field orthogonal to B0. This RF field is centered on the Larmor frequency of protons in the B0 field and causes the magnet moments to mutate their alignment away from B0 by some predetermined angle. The B1 field is produced in the imaging region of interest typically by an RF transmit coil that is driven by a computer-controlled RF transmitter with a RF power amplifier. During excitation, the nuclear spin system absorbs magnetic energy, and the magnetic moments precess around the direction of the main magnetic field. After excitation, the precessing magnetic moments will go through a process of free induction decay (FID), releasing their absorbed energy and returning to a steady state. During FID, NMR signals are detected by the use of a receive RF coil that is placed in the vicinity of the excited volume of a human body. The NMR signal is the secondary electrical voltage (or current) in the receive RF coil that has been induced by the precessing magnetic moments of the human tissue. The receive RF coil can be either the transmit coil itself or an independent receive-only RF coil. The NMR signal is used for producing MR images by using additional pulsed magnetic gradient fields that are generated by gradient coils integrated inside the main magnet system. The gradient fields are used to spatially encode the signals and selectively excite a specific volume of the human body. There are usually three sets of gradient coils in a standard MRI system that generate magnetic fields in the same direction of the main magnetic field and varying linearly in the imaging volume.
In MRI, it is desirable for the excitation and reception to be spatially uniform in the imaging volume for better image uniformity. In a standard MRI system, the best excitation field homogeneity is usually obtained by using a whole-body volume RF coil for transmission. The whole-body transmit coil is the largest RF coil in the system. A large coil, however, produces lower signal-to-noise ratio (SNR or S/N) if it is also used for reception, mainly because of its greater distance from the signal-generating tissues being imaged. Because a high signal-to-noise ratio is the most desirable in MRI, special-purpose coils are used for reception to enhance the S/N ratio from the volume of interest.
In practice, a well-designed specialty RF coil has the following functional properties: high S/N ratio, good uniformity, high unloaded quality factor (Q) of the resonance circuit, and high ratio of the unloaded to loaded Q factors. In addition, the coil device must be mechanically designed to facilitate patient handling and comfort, and to provide a protective barrier between the patient and the RF electronics. A further way to increase the SNR is to replace the single specialty coil with an array of smaller coils and through the use of multiple receivers add the signals together at the image construction stage. For this method to work effectively, the signals received from each coil in the array must collect or obtain signals from near the tissue and resonate with signals coupled from other coils. The coils must be decoupled to prevent signals from one coil interfering with signals from another coil.
One known method for decoupling coils includes overlapping adjacent coils by an amount necessary to cancel mutual inductance. However, this method is limited in that it requires the coils to be adjacent to each other, thus, limiting the ability to move the coils. Another known method for decoupling the coils includes canceling the mutual inductance by adding an extended loop to each coil, thus creating a transformer whose mutual inductance is designed to cancel the coupling between the two pickup coils. However, this method introduces capacitance between the coils and significant loss of Q factor. Still another known method for decoupling the coils includes using a capacitor common to two surface coils to cancel the mutual inductance between the surface coils. This series cancellation has a parallel equivalent where the series capacitor becomes a network. However, again, the coils must be adjacent to each other.
Thus, these known methods for decoupling coils restrict the design of coil arrangements (e.g., requiring coils to be adjacent to each other) in the MRI system. Therefore, the operation and control of these MRI systems is limited. For example, when designing a coil arrangement, only an overlap area that is acceptable for zero coupling may be used rather than a geometry that is optimum for imaging.